There are various different types of acquisition procedure in PET, including an emission scan, a transmission scan and a blank scan.
A typical emission scan begins with the injection of a solution including a tracer, which is a pharmaceutical compound including a radio-isotope with a short half-life, into the subject. The subject may be human or animal. The tracer moves to, and is typically taken up, in one or more organs in the subject according to biological and biochemical processes which occur within the subject. When the radio-isotope decays, it emits a positron, which travels a short distance before annihilating with an electron. This annihilation produces two high energy photons propagating in opposite directions. The PET scanner includes a photon detector array arranged (usually in a ring-shaped pattern) around the scanning area. If two photons are detected within a short timing window, a so-called “coincidence” is recorded along a line of response (LOR) connecting the two detectors. Coincidence counts along each LOR are incremented in a data storage part referred to as a sinogram bin each time a coincidence is detected between the corresponding detector pair. The output coincidence count data from each sinogram bin is typically processed using tracer uptake models and image processing techniques to obtain volumetric medical images and volumetric tracer uptake rate data for the subject.
It would be desirable to be able to ascertain detector efficiencies in order to be able to detect deterioration in the state of operation of detectors in the detector array, since upon deterioration of a detector, the scan results are accordingly reduced in accuracy, leading to poor image quality.
For quantitative results from PET images, attenuation correction forms one of the most crucial data correction stages. In a conventional PET scanner, the scanner is provided with one or more positron emitter rod sources, formed of a material such as 68Ge, which emit dual annihilation photons. To derive attenuation factors, two acquisitions using the rod sources are conventionally used—a blank scan, in which the subject being scanned is not present in the scanning area (typically, the scanner is empty except for the presence of the sources) and a transmission scan in which the subject is present in the scanning area. Since the source material is a positron emitter, the two photons arising from the annihilation of a positron and an electron are acquired in coincidence, in the same manner as with an emission scan. Conventionally, the results of the blank scan are then divided by the results of the transmission scan to derive an attenuation sinogram. The attenuation sinogram is then used to correct the emission scan for attenuation, although extra processing steps on blank scan data and transmission scan data are sometimes used.
Also, by collecting the coincidence count data during a blank scan and processing the count data using an efficiency estimation algorithm, individual detector efficiencies can be estimated. For non-rotating PET scanners, estimates of the detector efficiency may be calculated by adding together the counts detected between a particular detector and all the detectors that are in coincidence with it. This is known as a fansum and it is based on the following measurement model:Mij=εiεjAij  (1)
where Mij are the measured coincidence counts between detectors i and j, εi and εj are the intrinsic efficiencies of detectors i and j respectively and Aij are the ideal coincidence counts measured if the detectors had an ideal performance. Thus the measured counts Mij of a particular coincident pair of detectors are proportional to the product of the individual detector efficiencies. An estimate of each individual detector efficiency εi is made by summing the measured coincidence counts Mij over all detectors j in coincidence with detector i:
                              ɛ          i                =                              ∑            j                    ⁢                      M            ij                                              (        2        )            
The above equation is based on the assumption that:
                                          ∑            j                    ⁢                                    ɛ              j                        ⁢                          A              ij                                      =        constant                            (        3        )            
The assumption is reasonable providing that the efficiencies are not significantly different, that
  ∑  jAij is roughly independent of i and that the number of coincident detectors is sufficiently large. More accurate methods have been proposed which take into account noise and other variables. Further examples of efficiency estimation algorithms for calculating detector efficiencies in conventional PET scanners are described in “Maximum-Likelihood Estimation of Normalisation Factors for PET”, D. Hogg, K. Thielemans, T. Spinks and N. Spyrou, Nuclear Science Symposium and Medical Imaging Conference Record, IEEE, 2001.
In an alternative PET scanner arrangement, used for example in the ECAT EXACT3D™ PET scanner, a single photon source is provided for the transmission scan. In this scanner the source, a 137Cs pellet, is automatically moved by a transport mechanism having a spiral tube that is placed in front of the detectors when a transmission scan takes place. By forcing a liquid through the tube the source is moved at a constant speed through the scanner in a helical motion to provide a full 3D transmission scan.
As it uses a single photon source however, there is no second photon generated during an annihilation with which to form a coincident pair. With no coincident pair, no real coincidence counts are available from which a transmission sinogram could be created. The way this problem is solved is that the scanner in this mode measures only the detections furthest from the source, ignoring those registered in detectors close to the source, while tracking the position of the source. The source position is measured at four equidistant points about the helix. Since the speed of movement of the source is effectively uniform, its position can be calculated accurately to within one detector position over the entire length of the helix. From the position of the detector and the source, an artificial coincidence count is generated. The set of artificial coincidence counts detected during a transmission scan are collected and a transmission sinogram is calculated. However, detector efficiency estimates are not obtained from these results.
Some PET scanners have detector arrangements which rotate during an acquisition. Typically, the detectors are arranged in two or more banks which do not fully surround the subject. Alternatively, the detectors may be arranged in a non-ring-shaped pattern. In these cases, the model given by equation (1) above is no longer appropriate, as the rotation causes the direct inter-relation between detectors and entries in the sinogram bins to break down. Every sinogram bin typically contains counts detected by multiple detector-pairs. For some of these scanners, including the ECAT ART™ PET scanner, the transmission measurement is again performed using a single photon point source. However, the source is transported only axially relative to the (rotating) detector arrays. The ECAT ART scanner contains two such sources, one located on each of the opposing detector banks. The ECAT ART scanner is described in further detail in the article “The ECAT ART Scanner for Positron Emission Tomography: “1. Improvements in Performance Characteristics”, David W. Townsend et al., Clinical Positron Imaging, Vol. 2, No. 1, 1999 and the article “Design and Performance of a Single Photon Transmission Measurement for the ECAT ART” C. C. Watson, W. F. Jones, T. Brun, K. Baker, K. Vaigneur and J. Young, IEEE Medical Imaging Conference Record M9-02; 1998.
A phantom scan involves positioning a body, referred to as the phantom, containing a positron-emitting radioactive source material in the scanner. A phantom scan is either an emission or a transmission scan with a ‘phantom’ in the scanner instead of a patient. Typically, the ‘phantom’ is an object made out of plexiglass or suchlike and filled with water mixed with a radioactive substance. The phantom has known shape and attenuation characteristics. A phantom emission scan has in the past been used to calculate detector efficiency estimates. A problem with having to conduct a phantom scan is that the phantom must be handled by an operator to place the phantom inside the scanner and to subsequently remove the phantom. It is both inconvenient and time-consuming, since in any event an operator needs to be available to handle the phantom at the start and end of the procedure. It is also potentially hazardous, in terms of lifting the phantom and exposure to radioactivity.
It would be desirable to be able to estimate detector efficiency data in PET scanners of a type which use a single photon source for transmission scans, without the drawbacks of using a phantom scan.
In accordance with one aspect of the present invention there is provided a method of generating detector efficiency data for a positron emission tomography scanner including:
a detector array for generating detection data; and
a single photon source,
wherein the method comprises:
conducting an acquisition procedure using the single photon source to produce detection data; and
processing said detection data using an efficiency estimation algorithm to calculate data representative of the efficiencies of individual detectors in said array.
By enabling the estimation of detector efficiencies from an acquisition using the single photon source, detector efficiencies can be generated without significant inconvenience to an operator. The detector efficiencies may be derived from a blank scan acquisition conducted at the operator's convenience. Furthermore, the regular need for the use of a phantom scan procedure can be avoided.
In a preferred embodiment of the invention, detector efficiency estimates are made using the artificial coincidence counts generated during a blank scan acquisition made using the single photon source. Typically, artificial coincidence counts are the only suitable detection data made available as an output from a scanner of a type such as an ECAT EXACT3D PET scanner. However, detector efficiency data cannot be accurately estimated from artificial coincidence counts using known techniques, because the known measurement models do not apply. The present invention provides a new measurement model and exemplary efficiency estimation algorithms, which can be applied to artificial coincidence counts produced using blank scans.
Further features and advantages of the present invention will become apparent from the following description of preferred embodiments of the present invention, made by way of example only with reference to the accompanying drawings.